1. Field of the Invention
The present invention relates in general to split type or open superconducting magnets for magnetic resonance imaging (MRI), and in particular to such magnets having large diameter shield coils remotely spaced from the magnet poles.
2. Description of Prior Developments
Split type or xe2x80x9copenxe2x80x9d superconducting magnets are used in magnetic resonance imaging (MRI) scanners to produce magnetic fields required for patient imaging. Superconducting shield coils are typically used in each half of the split magnets to reduce stray electromagnetic fields. This type of shielding is referred to as active shielding.
A typical design of an actively shielded superconducting split type open MRI magnet assembly includes two generally cylindrical enclosures. A lower cylindrical enclosure and an upper cylindrical enclosure are interconnected by structural supports. Cryogenic and electrical connections are also provided between these enclosures or xe2x80x9chalvesxe2x80x9d.
The opposing cylindrical enclosures form two magnetic poles separated by a gap which contains an imaging region where a patient is imaged. Each enclosure contains several superconducting coils placed inside a liquid helium vessel. The helium vessel is located within a vacuum vessel and maintains the magnet at an operating temperature of about 4 K. The coils contained in the helium vessel are arranged in location as well as in magnetic polarity via the direction of the current they carry, so that they produce a substantially uniform field in a portion of the gap formed between the two enclosures while limiting the stray field outside the device to an acceptable level.
Each enclosure also includes one or more thermal shields as well as thermal insulation, located between the helium vessel and the vacuum vessel to keep heat leak to the 4 K environment within acceptable levels. The uniform internal field is produced by a main superconducting coil and several field shaping coils. In addition, one or more shield coils, are spaced away from the gap to reduce stray field released outside of the MRI scanner.
Sizing and positioning of the coils is done by those skilled in the art by using numerical codes based on static magnetic field equations, with the objective to achieve both the targeted field uniformity inside the imaging region, e.g. through minimizing terms in the Legendre polynomial series, and restricted stray field, e.g. through minimizing external field moments. Although the location and size of the coils cannot be defined arbitrarily, there is some freedom in positioning the coils, and different positions result in different amounts of conductor required for given uniformity and stray field requirements.
A significant advantage of a split type MRI magnet and scanner is the openness of the gap which is formed between the magnet poles. The open gap provides an enhanced view of the patient in the imaging region and allows medical personnel to directly access the patient as the patient is positioned within the gap.
Efforts are being made to increase the gap size to improve patient comfort, visibility and accessibility and to reduce the diameter of the magnet poles to further increase patient comfort, visibility and accessibility. Different magnet designs with different gap configurations can be produced within different envelopes yet provide the same field strength and field uniformity.
Such designs may require different aggregate amounts of superconducting material or xe2x80x9cconductorxe2x80x9d contained in the various coil designs. The conductor is usually the largest single cost item in an MRI magnet assembly, so it is desirable to minimize the amount of conductor required to produce a given field strength and uniformity.
Moreover, the designs with a larger conductor volume usually result in a higher peak field, Bpeak, and greater accompanying mechanical stress in the coils. These two parameters gradually increase with conductor volume until the risk of design failure becomes unacceptable. Peak field and stresses have a major impact on the progress of future open MRI designs with a target of high field exceeding 1T.
Accordingly, each split-type magnet design represents a compromise between its openness, which is related to gap size and pole diameter, its cost and its operating structural safety margins. Generally, for a given field strength, uniformity and stray field, the larger the maximum diameter of the magnet coils and/or the smaller the patient imaging gap, the lower is the amount of conductor required and the lower is the cost and structural risk of the magnet design. However, these potential cost reducing approaches can result in decreased patient comfort, visibility and accessibility.
Accordingly, a continuing need exists for a split type MRI magnet which provides improved openness through a large split and small diameter pole, yet which requires less superconducting material and which reduces the associated field and stress in a relatively lightweight superconducting magnet, while satisfying stringent uniformity and stray field requirements.
The present invention has been developed to fulfill the needs noted above and therefore has as an object the provision of an open or split type MRI magnet which provides improved openness or gap size, yet which reduces the total or aggregate amount of superconducting material required, and which also reduces the field and stress within the magnet coils under stringent requirements of uniformity and stray field.
A further object is to provide such a magnet with flattened or axially shortened main coils to reduce the amount of conductor material required and to reduce the amount of stress produced in the main coils.
Yet another object of the invention is to provide a split type MRI magnet with axially stepped field shaping coils to facilitate coil winding and to allow for the use of a single stepped coil support form.
Another object of the invention is to provide a split type MRI magnet with a series of field shaping coils located closely adjacent to the magnet poles and imaging gap.
Still a further object of the invention is to provide a split type MRI magnet having field shaping coils positioned radially outside of a recess formed in the magnet pole of a cryostat enclosure.
The above and additional objects are met by the present invention which is directed to an open split type superconducting magnet for an MRI scanner which uses superconducting shielding coils, with optional external iron shielding, to reduce stray magnetic fields. In accordance with the invention, the primary shield coil, i.e., that shield coil typically with the largest diameter having its current flowing opposite or negative with respect to the current in the main coil, can have a substantially greater axial distance from the imaging gap and advantageously has a substantially larger diameter than the main coil. The larger diameter of the shield coil does not affect the patient""s perception of openness as it has no effect on the patient""s field of view from the patient""s bed.
The main purpose of the negative primary shield coil is to compensate the external stray field produced outside of the magnet by the positive main coil and other coils. By doing so, the primary shield coil inevitably reduces the inner field in the imaging region, so the other coils have to grow in order to compensate this negative change. Contribution of the shield coil to the external field is determined by its magnetic moment, which grows as the square of its diameter. The shield coil of a larger diameter produces similar stray field with less ampere-turns and a reduction of the central field cancellation, hence it requires less conductor in other coils and in the whole magnet. More distant axial and radial positioning of the shield coil from the imaging gap additionally results in a smaller negative contribution in the imaging region, which further reduces the amount of conductor required for other coils.
This increased diameter of the primary shield coil can be carried out in either or both halves of the split magnet enclosure. Such a placement of the primary negative shield coil allows the main positive coil to have a smaller outer diameter while using the same amount of superconducting material, and while producing the same peak field and stress level in the conductor coils as in a main coil having a larger diameter. A smaller main coil can provide better gap openness which in turn provides for better patient comfort, visibility and accessibility.
As a result of increasing the outer diameter of the primary shield coil to a diameter greater than the outer diameter of the main coil, the outer shape or envelope of one or both halves of the magnet enclosures can be frustoconical, barrel shaped or xe2x80x9ctop hatxe2x80x9d shaped. Each of these shapes leads to a considerably smaller pole diameter at the expense of a larger central or top diameter on the respective half enclosure.
When a magnet enclosure is formed with a frustoconical shape or a barrel shape, the outer diameter of the pole formed by the cryostat enclosure is reduced and is positioned closer to the outer diameter of the main coil. One or both halves of the magnet enclosures may have an axial recess formed in the central portion of the magnet poles to accommodate gradient coils and/or RF coils. These design features provide better visibility and accessibility of the gap between the magnet poles and result in greater gap openness and patient comfort.
Accordingly, the present invention provides the same or greater gap openness while reducing the amount of conductor material required and while reducing the field and stress in a relatively lightweight superconducting magnet. To fully achieve these results, the outer diameter of the largest or xe2x80x9cprimaryxe2x80x9d shield coil is extended radially beyond the outer diameter of the main coil.
This placement of the primary shield coil reduces the amount of conductor material, coil field and coil stress compared with an open magnet having a conventional cylindrical enclosure and having a main coil with the same outer diameter. The conical surface of the vacuum vessel near the pole, enabled by and arising from such placement of the primary shield coil, also allows the outer diameter of the vacuum vessel pole to be reduced to a size closer to that of the main coil as compared with a magnet having a cylindrical enclosure, while keeping the necessary clearance between the 4 K helium vessel and the outer vacuum vessel the same.
The openness of the gap between the poles is not seriously affected by an enlarged shield coil because of the axial spacing of the shield coil remotely away from the gap which retains the same patient""s field of view. The comparison noted above assumes that the outer diameter of the main coil is kept the same as that of a conventional cylindrical magnet enclosure, but that the amount of conductor material (and cost of material) is reduced. However, considering the trade off between the amount of conductor material required and the pole diameter, a magnet designer can now reduce the outer diameter of the main coil while keeping the amount of coil material the same. This results in a more open gap and a magnet having a similar cost and reliability as conventional cylindrical designs.
Another feature of the invention is the placement of one and preferably two or more magnetic field shaping coils at a position closely axially adjacent to a magnet pole and imaging gap. The closer the field shaping coils are located to the magnet pole, the less superconducting coil material is required to produce the desired effect.
In one embodiment, the field shaping coils are arranged in substantial radial alignment on an inner wall of the magnet enclosure so that the axial spacing of these coils from the imaging gap is minimized. This radial coplanar or xe2x80x9cplanetaryxe2x80x9d arrangement of coils reduces the amount of conductor material required as compared to field shaping coils located axially further away from the magnet pole and imaging gap.
However, such planetary positioning of coils can pose coil winding problems and may necessitate a slight axial staggering of these field shaping coils to facilitate coil winding. This can be accomplished by using axially and radially stepped winding bobbins or mandrels. Axial spacing of the field shaping coils from the gap should be minimized to reduce the amount of conductor material required.
It is possible to separately wind the coils on individual bobbins and then mount them in machined grooves in the magnet enclosure, and in particular, in the inner wall of the helium vessel mounted within the enclosure. It is also possible to form bobbin or coil support structures within the enclosure, such as on the inner wall of the helium vessel, and wind the coils in place on the helium vessel.
A particularly advantageous design positions the main coil and primary shield coil as noted above, and further positions the field shaping coils radially inwardly of the main coil and axially inwardly of the primary shield coil. Preferably, the field shaping coils are positioned axially between the main field coil and the primary shield coil.
The aforementioned objects, features and advantages of the invention will, in part, be pointed out with particularity, and will, in part, become obvious from the following more detailed description of the invention, taken in conjunction with the accompanying drawings, which form an integral part thereof.